KR20220019718A - Regulation of electron transport kinetics in E-DNA type sensors - Google Patents

Abstract

An improved electrochemical sensor in which a recognition element is deposited with a secondary charge transfer control moiety, eg, an oligonucleotide. The secondary moiety modulates the electron transport kinetics to enhance the frequency dependence of the sensor gain, enabling the use of dynamic differential drift correction techniques and measurements requiring target-insensitive signal drift in parallel with the target-dependent output. The secondary moiety also increases the sensor's gain and signal-to-noise and can be used to enable uncorrected measurements. The use of an accurate drift-corrected in vivo sensor with multiple measurements of analyte concentration per minute in flowing blood is described.

Description

Regulation of electron transport kinetics in E-DNA type sensors

CROSS-REFERENCE TO RELATED APPLICATIONS: This application claims the benefit of priority to U.S. Provisional Application Serial No. 62/860,772, entitled "Modulating Electron Transfer Kinetics in E-DNA-type Sensors," filed on June 12, 2019; , the contents of which are incorporated herein by reference.

REFERENCE TO SEQUENCE LISTING, TABLE, OR COMPUTER PROGRAM LISTING COMPACT DISK APPENDIX: This application contains a sequence listing electronically submitted in ASCII format, which is hereby incorporated by reference in its entirety. The ASCII copy created on June 12, 2020 is named UCSB029PCT_SL.txt and is 4,207 bytes in size.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT: This invention was made with government support under Grant No. R01 EB022015 awarded by the National Institutes of Health.

A major goal in the field of biosensing is the continuous (or at least high-frequency) real-time measurement of target analytes in complex samples and environments, such as those found in vivo. Achieving this capability for a wide range of target analytes could open up numerous applications in research, environmental monitoring, industry and medicine. Reagent-free electrochemical sensors contain a recognition element such as an aptamer or protein that selectively binds to a target molecule, and detection is difficult if this binding induces a change in the electron transport kinetics between the recognition element of the sensor and another element. possible. Such sensors will be referred to herein broadly as “receptor-based sensors,” which include several platforms including DNA-based and aptamer-based electrochemical sensors (“E-DNA” and “E-AB” sensors, respectively). do. As with most sensors, reagent-free electrochemical sensors tend to drift over time, especially when the sensor is exposed to complex samples such as whole blood or other bodily fluids, where the sensor output independently changes the target analyte concentration. do.

In some cases, sensor drift may be corrected using a drift correction method, such as a kinematic differential measurement (KDM) technique. KDM is described, for example, in Ferguson et al., Real-Time, Aptamer-Based Tracking of Circulating Therapeutic Agents in Living Animals, Science Translational Medicine 27 Nov 2013, Vol. 5, Issue 213, pp. As described in 213ra165, it self-corrects for signal drift and improves the signal-to-noise ratio. This technique exploits the square wave frequency dependence of the signal in this kind of sensor. In particular, electron transfer is faster in target-bound receptors than in untargeted receptors. When square wave voltammetry is performed at a higher frequency and vice versa at a lower frequency, this difference in kinetics results in an increase or decrease in the coupling induced current. Conveniently, these two outputs drift together, so the difference is used to effectively compensate for baseline drift.

For appropriate recognition factors, KDM provides a convenient way to compensate for drift. However, the implementation of KDM depends on whether there is a strong dependence between the signal behavior and the frequency of electrode interrogation. In particular, for successful implementation of KDM, there must be at least one response frequency at which target binding significantly alters the signal, there must be at least one non-reactive or minimal response or negative response frequency, and the sensor output must be completely or They are usually independent or even decrease in response to goals. That is, at the selected response frequency, the gain should be high to impart sensitivity and/or high dynamic range and at the selected non-response frequency, the gain enables a similar correction adjustment to capture the sensor signal due to KDM or drift. must be at least, 0, or negative to For example, for several aptamer recognition elements, target binding produces signal-on behavior at high frequencies and signal-off or no-response outputs at low frequencies.

However, some recognition elements, e.g. some aptamers, are configured such that the signal output is not a strong function of frequency, e.g., target binding elicits a signal-on response over a standard range of SWV frequencies. For example, when an E-AB sensor using an aptamer that binds camptothecin is used as a recognition element for detection of the chemotherapeutic agent irinotecan, the sensor has a significant gain at all frequencies. Since there is no low-gain frequency range, there is no reference point to account for the sensor's natural drift over time. For these frequency-insensitive recognition elements, KDM may not be implemented in standard sensor configurations. Therefore, there is a need for novel sensor configurations and measurement techniques that enable drift correction in receptor-based sensors that cannot follow this approach.

Another problem with using receptor-based sensors is that they require calibration. The raw absolute output signal generated by receptor-based sensors varies significantly from sensor to sensor due to changes in manufacturing methods. This variability is likely due to the difference in the microscopic surface area of the sensor electrode and the difference in the packing density and the total number of redox reporters exchanging electrons. Likewise, non-specific binding by non-target species present in a sample often causes variability in sensor output. In a controlled situation, this problem can be circumvented through a calibration step in which the output is measured when the sensor is exposed to a sample of a known target concentration. Such calibration may be sufficient for many applications, such as, for example, continuous monitoring of an exogenously applied target or benchtop applications where blank or standard samples can be readily used. However, important calibration procedures cannot be applied in many settings, for example in the case of in vivo monitoring. For ex vivo applications, such as point-of-care applications, the need for calibration increases the complexity, time, and cost of the method and in many cases the calibration step is beyond the practical or economical limits of clinical or bedside applications. Therefore, there is a need in the art for a calibration-free measurement method that can accurately quantify the concentration of a given molecular analyte regardless of sensor-to-sensor variation or sensor drift.

The inventors of the present invention have advantageously found a way to modulate electron transport kinetics in receptor-based sensors using secondary moieties deposited with recognition elements. The secondary moiety did not respond to the presence of the target, but modulated the frequency dependence of electron transport by the sensor system. It can do so by transferring the electrons themselves. Or it may do so by modulating the electron transport behavior of the target binding recognition element. By including these molecules together with the recognition element, the electron transport kinetics between the recognition element and the electrode can be modulated to a beneficial effect. First, the gain of the sensor may be increased. Second, the frequency dependence of the sensor gain can be improved.

In a first aspect, the scope of the present invention encompasses novel sensor designs wherein the recognition element of a receptor binding sensor is deposited with a secondary charge transfer-modulating moiety on the sensing electrode. In some embodiments, the sensor is a DNA type sensor, such as an aptamer based sensor. In some embodiments, the secondary charge transfer-regulating moiety is a redox-reporter labeled oligonucleotide, such as a short oligonucleotide of 2 to 20 bases. In some embodiments, the secondary charge transfer-regulating moiety is an oligonucleotide, such as a short oligonucleotide of 2 to 20 bases without a redox-reporter.

In another aspect, the scope of the present invention includes the use of the novel sensor of the present invention for drift correction. The secondary charge transfer-regulating moiety provides a reference signal that is insensitive to the target concentration and drifts in parallel with the measurement signal generated by the recognition element. Thus, the drift effect can be quantified and subtracted from the measured signal to provide a drift-corrected measure.

In another aspect, the gain of the electrochemical sensor is enhanced by the use of a sensing element comprising a recognition element deposited with a secondary charge transfer-modulating moiety. The improved gain enables sensor operation in challenging sample types, such as, for example, whole blood, which is flowing whole blood in vivo, where a high signal-to-noise ratio is required to analyze the target concentration.

In another aspect, the novel sensors of the present invention can be used in uncalibrated operation. With co-deposited secondary charge transfer-regulating moieties, a stronger frequency dependence of the gain can be created. Using the frequency dependence of the gain, a normalized sensor output corresponding to the sensor-to-sensor variation affecting the absolute output can be obtained. The ability to utilize the electrochemical sensors of the present invention in calibration-free operation allows for extended use in vivo or in other situations where calibration is cumbersome or impractical.

In another aspect, the scope of the present invention includes novel camptothecin binding aptamers. The novel aptamers confer higher gain than the camptothecin aptamers of the prior art when used in an electrochemical sensing platform.

Included in the context of the present invention.

1a, 1b, 1c and 1d. 1A is a diagram illustrating an exemplary prior art electrochemical aptamer based (E-AB) sensor. The figure depicts unbound aptamers 103 and 107 modified with a redox reporter covalently attached to the gold electrode 101 via an alkane-thiol self-assembling monolayer 102 . In the absence of a specific target, the aptamer is partially or fully unfolded (103). When the target molecule 104 binds to the aptamer 105, the conformational change alters the proximity of the redox reporter 108 and electron transfer to the electrode observed in voltammetry (eg, square wave) interrogation. Improve the efficiency of the signal (star). Figure 1B: In the finished sensor, a 75 mm gold working electrode was bundled with a platinum counter electrode of the same diameter and a silver/silver chloride reference electrode, sufficient for placement through a 22-gauge guide catheter into one of the external jugular veins of a live rat. It manufactures small and flexible devices. 1C shows sensor placement in a rat 101 , with the sensor bundle 102 external to the animal and the sensing electrode 103 placed in the jugular vein 104 . 1D shows the different signal outputs observed in the absence and presence of a target.
Figures 2a, 2b, 2c and 2d. Figure 1A: The parent aptamer is expected to fold into a G-quartet that is thought to be the target-binding site flanked by a 12 base pair stem. Figure 2b: Using the intrinsic fluorescence of irinotecan when free in solution, the aptamer exhibits a dissociation constant of 475 nM. Figure 2c: However, when the aptamer is redox-reporter-modified and immobilized on the interrogation electrode of the sensor, especially when placed in undiluted whole blood, its affinity and signal gain are significantly reduced, and this binding curve is A square wave frequency was used. 1D: Nevertheless, the E-AB sensor responds rapidly when challenged with irinotecan (here in buffer). This coupling curve used a square wave frequency of 500 Hz and a repetition rate of 0.2 Hz.
3a, 3b and 3c. Figure 3A: The parent aptamer destabilizes the stem loop of the aptamer either through introduction of one (CA40_1MM) or two (CA40_2MM) mismatches or through cleavage of the stem (CA36, CA32, CA28, CA16), resulting in higher gain E -redesigned to generate an AB signal (thus increasing the population of unfolded molecules maintained in equilibrium to respond to the target). Figure 3b: When challenged in a simple buffer solution, all redesigned variants showed higher gain than the parent aptamer, and the most unstable ones (CA40_2MM, CA32, CA28, CA16) showed the greatest signal gain. Figure 3c: When tested in whole blood, the gain and affinity decrease but nevertheless the best performance still supports high gain E-AB detection.
4a, 4b and 4c. 4A is a diagram showing a sensor of the present invention. The sensor comprises a gold electrode 101 functionalized with a SAM layer 102 , wherein target-binding aptamers 103 , 107 modified with a redox reporter by the SAM layer 102 are secondary moieties, in this case It is deposited with a redox reporter-modified oligonucleotide strand (105). Unbound aptamer 103 is shown. When the aptamer 107 binds to the target molecule 108, it undergoes a conformational change, where the proximity of the redox reporter 109 to the electrode 101 changes the electron transport kinetics and the resulting signal. Figure 4b: The signal gain (relative signal change between no target and saturation target - ie 100 mM) of the original E-AB sensor (100:0 curve) is a relatively small function of the square wave frequency. As the amount of linear strands increases (up to a maximum ratio of 50:50), an increasingly strong frequency dependence is observed, despite a corresponding decrease in the maximum gain. Figure 4c shows KDM drift correction (difference in relative signals visible at 10 and 120 Hz) for the target in the clinically relevant range (0.5 mM to 15 mM) using a 50:50 mixture of both strands in both buffer and undiluted whole blood. Describe the sensor fabricated using
5a and 5b. 5A: Sensor output over time in response to an intravenous injection of irinotecan (60 mg/kg ¹ ) from time 20 min to ˜55 min. 5B: Close-up of the excretory phase from ˜55 min to 86 min. The higher peak concentrations reached in this experiment lead to longer measurement runs, which in turn improve the accuracy of estimates of relevant pharmacokinetic parameters.
6a, 6b and 6c. Figure 6a shows the signals collected by the sensor of the present invention at high (120 Hz) and low (10 Hz) frequencies. Both signals drift significantly, but at the same time. Figure 6b: Taking the difference (KDM) between normalized high (120 Hz) and low (10 Hz) frequency signals produces a stable baseline. 6C: Control injection of saline “blank” or a second chemotherapeutic agent (5-fluorouracil, often co-administered with irinotecan) does not produce any measurable sensor response.
Figure 7. Figure 7 depicts irinotecan levels measured after multiple intravenous injections in live mice. The black line indicates the fit of each injection dataset to the two-compartment pharmacokinetic model.
8a and 8b. 8A and 8B depict E-AB sensor output collected in vitro in buffer (FIG. 8A) or flowing whole blood (FIG. 8B). A K _D = 16.0±1.3 μM is observed for the KDM signal in the buffer. In whole blood, K _D rises to 191.2±23.4 μM. Error bars correspond to the standard deviations observed for three independently fabricated and tested sensors.
9a, 9b, 9c, 9d, 9e and 9f. 9a, 9b, 9c, 9d, 9e and 9f are 100% C28 mutant aptamers (SEQ ID NO: 6) functionalized with -SH at the 5' end for attachment to the SAM functionalized gold electrode surface and 3 with methylene blue. ' depicts the effect of frequency on signal gain for a mixture of a C28 aptamer functionalized at the end and an oligonucleotide secondary moiety. The signals shown are the average of three sensors of 100 μM irinotecan dissolved in DMSO dissolved in PBS at room temperature; E _increase = 0.001V; Amplitude = 0.05V. 9A, 9C, 9D, 9E and 9F depict signals from pure CA28 (“100:0”) and 50-50% mixtures of C28 with oligonucleotides (“50:50”). Figure 9a: 5nt oligonucleotide secondary moiety, SEQ ID NO: 9. Figure 9b shows pure CA28 (“100:0”) and various mixtures of C28 and 10nt oligonucleotide secondary moieties, SEQ ID NO: 8 (50 :50-50% CA29 and 50% secondary moieties 40:60-40% CA28 and 60% secondary moieties, 20:80-20% CA28 and 80% secondary moieties). Figure 9c: 15nt oligonucleotide secondary moiety, SEQ ID NO: 10. Figure 9d: 30nt oligonucleotide secondary moiety, SEQ ID NO: 11. Figure 9e: 40nt oligonucleotide secondary moiety, SEQ ID NO: 12 Figure 9f: 60nt oligonucleotide secondary moiety, SEQ ID NO: 13.

The scope of the present invention includes novel electrochemical sensor designs and methods of use thereof. The present invention relates in particular to electrochemical sensors, which may be collectively referred to as “receptor-based electrochemical sensors”. Receptor-based electrochemical sensors include a recognition element functionalized with a redox reporter, the recognition element selectively binding to a target molecule. The redox reporter functionalized recognition element is coupled to or otherwise associated with the electrode. When energized by application of voltage and/or current pulses, the redox reporter exchanges electrons with electrodes, producing a measurable signal. Binding of a target molecule to a recognition element modulates the kinetics of this electron transport and the resulting change in signal correlates with target concentration.

The electrochemical sensor of the present invention includes a secondary charge transfer-modulating moiety that modulates the signal from the recognition element and provides various advantages. First, the secondary moiety can enhance the gain of the sensor. Second, the secondary moiety imparts or enhances the frequency dependence of the sensor output signal, enabling drift compensation methods and enabling uncompensated operation of the sensor. Various elements and embodiments of the invention are described below.

electrochemical sensor . The improved sensor design of the present invention can be applied to any electrochemical sensor. In a first embodiment, the sensor of the invention will be an E-DNA type sensor that uses a target-binding polynucleotide such as an aptamer as a recognition element. Any E-DNA sensor design or configuration known in the art may be used. In some embodiments, the E-DNA sensor is an E-AB sensor, wherein the recognition element comprises an aptamer known in the art. Polynucleotide recognition elements can include DNA, RNA, or non-natural nucleic acids as well as hybrids of the foregoing.

In the E-DNA sensor, one or more selected portions of the working electrode are functionalized with polynucleotides. The polynucleotide may be conjugated to or otherwise associated with the electrode surface by any suitable chemistry, for example, covalent bonding, chemisorption or adsorption. Alkane thiol monolayers can be used to bond polynucleotides to electrode surfaces, particularly suitable for gold electrode surfaces.

The scope of the present invention is not limited to E-AB sensors. The scope of the present invention further includes any electrochemical sensor wherein target binding to the recognition element produces a measurable change in the rate of electron transport detectable by the sensing element. In various embodiments, the recognition element is not a nucleic acid, and sensors that use, for example, proteins (eg, antibodies or fragments thereof), chemical species, and other molecules as recognition elements, are also within the scope of the present invention.

In one embodiment, sensors that use unmediated electrochemical sensing including the use of direct electron transfer and redox equilibrium as a way to generate a signal can be used. Such biochemical sensors include, for example, sensors that detect consumption or production of species belonging to redox couples. Other sensors may use mediated electrochemical analysis, ie, use a redox species mediator for electron transfer and establish a redox equilibrium. Examples of mediated and non-mediated sensors are described in Sander et al. 2015, A Review of Nonmediated and Mediated Approaches. Environ. Sci. Technol . 49:5862-5878.

Additional sensor types include chemically modified electrodes, immune sensors, oligopeptide-based sensors, and enzyme sensors.

Other sensor types that may be used include ligands, for example, Prasad, 2004, Prasad, 2004, The Role of Ligand Displacement in Sm(II)-HMPA-Based Reductions. J Am Chem Soc. 2004;126(22):6891-4 is a sensor based on electron transport changes due to targeted binding-induced displacement of hexamethylphosphoramide and samarium(II) iodide.

Another exemplary sensor type is the reconstitution energy of a redox reporter, eg, Plumb, 2003, Interaction of a Ferrocenoyl-Modified Peptide with Papain: Toward Protein-Sensitive Electrochemical Probes. Bioconj Chem. 2003;14(3):601-6. doi: ferrocenoyl-peptide described in 10.1021/bc0256446; or see, for example, Feld 2012, Trinuclear Ruthenium Clusters as Bivalent Electrochemical Probes for Ligand-Receptor Binding Interactions. Langmuir. 2012;28(1):939-49. It is based on the transformation of the trinuclear ruthenium cluster described in doi: 10.1021/la202882k.

Another sensor type that may be used is that a scaffold-attached redox reporter is, for example, Ge 2010, A Robust Electronic Switch Made of Immobilized Duplex/Quadruplex DNA. Angew Chem Int Ed. 2010;49(51):9965-. Doi: Double DNA, Quadruple DNA and DNA Nanoswitches as described in doi: 10.1002/anie.201004946 or Cash 2009, An Electrochemical Sensor for the Detection of Protein-Small Molecule Interactions Directly in Serum and Other Complex Matrices. J Am Chem Soc. 2009;131(20):6955-7. It is a sensor based on a sterically induced change in the efficiency of accessing primary electrode surfaces, such as the DNA-containing small molecule recognition element described in doi: 10.1021/ja9011595.

In some embodiments, the recognition element of the sensor comprises a polypeptide. For example, in one embodiment, the polypeptide may comprise a receptor or ligand-binding domain thereof, wherein the target species is a ligand. In one embodiment, the polypeptide is an antibody or antigen-binding fragment thereof, wherein the target species is an antigen.

In one embodiment, the electrochemical sensor is an E-AB sensor comprising a double stranded sensor, wherein the redox species are on separate strands, a portion of which is complementary or otherwise reversible binding to a portion of the aptamer. can In the presence of the target species, a strand of the redox species is liberated from the aptamer, causing the target species to bind to the aptamer and the redox species contacting or proximate to the electrode.

In some implementations, the electrochemical sensor is a signal-on type sensor such that target binding enhances a signal, and in other implementations the electrochemical sensor may include a signal-off configuration, wherein target binding reduces the output signal or Remove .

Redox reporter. In a first embodiment, each recognition element of the electrochemical sensor is functionalized with one or more redox reporters. Binding of the target species causes the recognition element to change its composition, causing the location (or accessibility to the electrode) of one or more redox reporters to be appreciably altered. Similarly, in many embodiments of the invention, the secondary charge-controlling moiety is modified with one or more redox reporters. The redox reporter of the recognition element and/or secondary moiety can comprise any composition of matter that interacts with the electrode, such that changes in accessibility or proximity to the electrode cause a change in electron transport kinetics. Exemplary redox species include methylene blue, ferrocene, viologen, anthraquinone or any other quinone, ethidium bromide, daunomycin, organo-metal redox label such as a porphyrin complex or crown ether cycle or linear ether , ruthenium, bis-pyridine, tris-pyridine, bis-imidazole, ethylenetetraacetic acid-metal complex, cytochrome c, plastocyanin and cytochrome c'.

sensor component. The electrochemical sensor may include one or more working electrodes to which a plurality of recognition elements are coupled. Exemplary recognition element densities may range from 1x10 ¹⁰ to 1x10 ¹³ molecules/cm ² . The working electrode may be, for example, any metal surface that forms a bond with a thiol or amine; gold; any gold coated metal (titanium, tungsten, platinum, carbon, aluminum, copper, etc.); any suitable electrode material for electrochemical sensing, including bare palladium electrodes, carbon electrodes, and the like.

The working electrode may be configured in any desired shape or size. For example, paddle-shaped electrodes, rectangular electrodes, wire electrodes, electrode arrays, screen printed electrodes, and other configurations may be used. For in vivo measurements, a thin wire configuration is advantageous because the low-profile wire can be inserted into a vein, artery, tissue or organ and does not interfere with blood flow in the blood vessel or cause significant damage to the tissue. For example, wires of 1-500 μm in diameter, such as 50 μm, 75 μm, or 100 μm, can be used.

The electrochemical sensing system of the present invention further comprises an auxiliary electrode or counter electrode, for example a platinum auxiliary electrode. The electrochemical sensing element may be used in conjunction with a reference electrode, such as an Ag/AgCl electrode, or other reference electrode known in the art. The electrochemical sensor of the present invention may be configured as a two-electrode or three-electrode system, and is suitably configured for performing chronoamperometric measurements. The electrode containing cell system may include a mixing chamber or other vessel in which the electrode is present and in contact with the sample.

The sensor and electrode system may include an assembly for obtaining a paradic current measurement when placed on or exposed to the sample. The assembly may include a housing. For example, for placement in the body of a living organism, the housing may include a needle, catheter, or other implantable structure. For ex vivo applications, the housing may contain wells, microfluidic containers, or other structures such as those found in lab-on-chip devices.

The sensing element of the present invention may be functionally coupled to an appropriate component for performing electron transport measurements. A measurement component may include two or more devices in electrical and/or network connection to each other or may include a single integrated device. A first component for performing a measurement includes a device or combination of devices capable of delivering excitation voltage pulses of a desired magnitude, frequency, and waveform to the sensing element. The measurement component may include a potentiometer or other voltage source and voltage controller for imposing a voltage step on the working electrode.

A second component for performing the measurement includes a device or combination of devices capable of obtaining a time-resolved current output from the sensing element. Such components will include circuitry for reading sensor outputs and storing such outputs or storing such outputs or routing the outputs to other devices including the components as analog-to-digital converters, amplifiers and digital storage media. .

target species. The sensor of the present invention relates to the detection of a target species. A target species may include any inorganic or organic molecule, for example: a small molecule drug, metabolite, hormone, peptide, protein, carbohydrate, nucleic acid, lipid, hormone, metabolite, growth factor, neurotransmitter, or nutrient. there is. The target may contain contaminants or contaminants. The target may include a toxin. Targets may include pathogen-derived or pathogen-derived factors, or viruses or cells. In some embodiments, target species include drugs with significant side effects, such as chemotherapeutic drugs or drugs with a narrow therapeutic index, where accurate measurement of blood levels is important to ensure safe administration or minimal side effects.

secondary moiety. The sensing element of the present invention comprises a charge transfer-regulating moiety, referred to herein as a secondary moiety, which is a non-target binding molecule co-deposited or otherwise present on the sensing element. In one embodiment this secondary moiety comprises a redox reporter that (1) transfers electrons faster or slower than the recognition element and (2) is not responsive to the presence of the target.

In a second embodiment, the secondary moiety does not contain its own redox reporter. Instead, it interacts with a reporter or recognition element to modulate the electron transport rate of the redox moiety on the recognition element. These interactions may be driven by steric or electrostatic, or more specific interactions such as hydrogen bonding or hydrophobic interactions.

For convenience of reference, the charge transfer control moiety may be referred to herein as a “secondary moiety”. In a first embodiment, it will be a single species of molecule wherein a plurality of such molecules are mixed with a plurality of recognition elements on the electrode surface. However, the term "secondary moiety" will be understood to include one or more different types of molecules, e.g., in some embodiments, mixtures of two or more different types of molecules, i.e. mixtures that are heterogeneous in any proportion. .

One mechanism by which this regulatory secondary moiety acts is to interfere with electron transfer between the redox reporter on the recognition element and the working electrode. This interference generally slows the rate of electron transfer between the redox reporter on the recognition element and the electrode, where the rate of electron transfer to the bound and unbound states of the recognition element responds differentially to the interference generated by the secondary moiety. do. Thus, the use of a modulating moiety can alter the frequency dependence of the target induced current generated during SWV or other excitation sequences. It is also possible to improve the propagation rate separation between the engaged and uncoupled states of the recognition element, thereby improving the sensor gain.

The secondary moiety can be any species that can be deposited with recognition elements, including biomolecules such as oligonucleotides and proteins, or other compositions comprising polymeric organic materials or other chemical entities. In a first embodiment, the secondary moiety comprises an oligonucleotide. Advantageously, for E-AB and other E-DNA type sensors, the use of oligonucleotides as regulatory moieties allows for single attachment chemistries (eg, thiolated 5' DNA ends for attachment to gold or other surfaces). enables the simultaneous deposition of aptamers and secondary moieties in a single step by In a primary embodiment, the oligonucleotide includes DNA, but the scope of the invention includes other nucleic acids, including non-naturally occurring nucleic acid analogs.

Oligonucleotides can be of any length, for example 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, 20, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, 35, 36, 37, 38, 39, 40, 41, 42, 43, 44, 45, 47, 48, 49, 50, 50-55, 55-60, 60-65, 65-70, 70-75, 75-80, 80-85, 85-90, 90-95- 95-100, 100- 110, 110-120, 120-130, 130-140, 140-150, 150-160, 160-170, 170-180, 180-190, 190-200, or more than 200 nucleotides. In general, the longer the oligonucleotide, the greater the interference effect on the charge transfer rate between the redox-modified recognition element and the working electrode. In some embodiments, the oligonucleotide comprises 5-10, 5-20, 10-15, 10-20, or 10-50 nucleotides.

In one embodiment, the oligonucleotide comprises a sequence of thymine and adenine. In one embodiment, the secondary moiety is an oligonucleotide consisting solely of the base thymine. In one embodiment, the secondary moiety is an oligonucleotide composed solely of adenine. In various embodiments, the nucleotides comprise at least 10%, at least 20%, at least 30%, at least 40%, at least 50%, at least 60%, at least 70%, at least 80%, at least 90% or at least 95% thymine. . In various embodiments, the nucleotides comprise at least 10%, at least 20%, at least 30%, at least 40%, at least 50%, at least 60%, at least 70%, at least 80%, at least 90% or at least 95% adenine. .

In one embodiment, the oligonucleotide is an oligonucleotide selected from the group consisting of SEQ ID NO: 8, SEQ ID NO: 9, SEQ ID NO: 10, SEQ ID NO: 11, SEQ ID NO: 12 and SEQ ID NO: 13 or at least 80% for an oligonucleotide selected from the group consisting of SEQ ID NO: 8, SEQ ID NO: 9, SEQ ID NO: 10, SEQ ID NO: 11, SEQ ID NO: 12 and SEQ ID NO: 13, oligonucleotides having at least 85%, at least 90%, at least 95%, or at least 99% sequence identity. In various embodiments, the oligonucleotide is a subsequence of any of the foregoing.

In an alternative embodiment, the secondary moiety comprises a polypeptide. For example, the secondary moiety comprises a polypeptide of 2 to 1,000 amino acids, for example, a polypeptide of 2-5, 5-10, 10-15, 15-20, 20-25, or 30 or more amino acids. may include Exemplary polypeptides include flexible and essentially disordered sequences, such as chains of threonine, serine, proline, glycine, aspartic acid, lysine, glutamine, asparagine or alanine, or mixtures thereof. It may comprise naturally occurring proteins that are essentially disordered. It may include folded proteins.

In an alternative embodiment, the secondary moiety comprises an organic composition. In one embodiment, the organic composition comprises a polymer such as polyglycol, polyacid, polyalcohol, polyether, polyester, polyvinyl, or polysaccharide.

In a first embodiment, the secondary moiety is modified with one or more redox reporters. In one embodiment, the secondary moiety is modified with a single redox reporter. For example, in the case of a secondary moiety comprising an oligonucleotide, the redox reporter may comprise a 3' attached moiety, eg, methylene blue. The redox reporter of the secondary moiety will generally be the same chemical entity as the recognition element is modified, but in an alternative embodiment, the secondary moiety may comprise a redox reporter different from the redox reporter of the recognition element. .

In an alternative embodiment, the secondary moiety does not include a redox label. In this embodiment, the electron transport-modulating properties of the secondary moiety derive entirely from the effect of the moiety on the charge transport kinetics of the redox-reporter-modified recognition element.

The absolute abundance of the charge transfer-regulating moiety can be anything sufficient for effective signal transduction properties, for example, at densities of 1x10 ¹⁰ to 1x10 ¹³ molecules/cm ² . The relative ratio of the modulating portion to the recognition element of the sensor will affect the degree of signal speed modulation. The percentage of secondary moieties (i.e., the ratio of secondary moieties present to the total recognition element to secondary moieties present) may vary, for example, a regulatory moiety of about 5%, about 10% , about 20%, about 30%, about 35%, about 40%, about 45%, about 50%, about 55%, about 60%, about 70%, about 80%, or about 90%. In one embodiment, the percentage of regulatory moieties is between 40 and 60%. In one embodiment, the percentage of regulatory moieties is about 50%. This ratio can be controlled by controlling the concentration of the two species during manufacture or by changing the deposition time.

How to use. The scope of the present invention includes methods of using the novel sensors of the present invention. A general electrochemical sensor operation is described below, followed by a description of specific uses in which the novel sensing element of the present invention may be used.

pulsed voltammetry. In a first embodiment, the electrochemical sensor may be used with a pulsed voltammetric method. Pulsed voltammetry is an electrochemical measurement technique in which a functionalized electrode of a sensor is subjected to a series of potentials applied to a pulse waveform and one or more measurements of the system current are evaluated in each cycle. In the practice of the present invention, a series of potential pulses will include one or more suitable potential pulses, which are applied potentials at a value that produces a measurable current output from the redox reporter species used in the system. In general, the series of potential pulses will include pulses at, in the vicinity of, or comprising (i.e., in a swept voltage method) the redox potential(s) of one or more redox reporter species of the sensing element, i.e., the sensor It is possible to stimulate and induce Faraday current flow between the recognition element and the redox reporter of the secondary moiety and the electrode. For example, electrical excitation of the redox reporter may induce a transient current flow between the redox reporter and the electrode substrate (or between the electrode substrate and the redox reporter depending on the system configuration). When stepping the voltage of a working electrode, such as the substrate of an E-AB sensor, the working electrode becomes either a stronger reducing agent (if stepping to a more negative potential) or a stronger oxidizer (if stepping to a more positive potential in this case). do. Within an appropriate range (eg, near or above the redox potential of the redox reporter), this voltage step will induce a Faraday current flow between the redox reporter of the sensing element and the electrode substrate. When an electric current flows, the pool of electrons mobilized by the excitation is depleted and the current decays with a rate and amplitude that depends on the binding state of the recognition element of the sensor, where target binding, for example, reduces the proximity of the redox reporter to the working electrode. By changing it, we induce a faster or slower transfer of current. Thus, the ratio of target-bound recognition element to untargeted recognition element, which is proportional to the concentration of the target species in the sample, will determine the current decay rate and amplitude observed for the sensor as a whole.

As used herein, “current” will refer to the flow of electrons measured by a sensor disposed on a sample. For example, the current may comprise a flow of electrons from a redox reporter to an electrode, or may comprise a flow of electrons from an electrode to a redox reporter.

In a first embodiment, pulsed voltammetry includes square wave voltammetry (SWV) as is known in the art. In SWV, a square voltage waveform is applied to the working electrode and the cell's paradic current is measured. The current may be resampled twice during each square wave cycle, once at the end of the forward pulse, and again at the end of the reverse pulse. This technique determines the charging current by delaying the current measurement to the end of the pulse. The current difference between the two measurements is plotted against the applied potential.

The scope of the present invention further includes other pulsed voltammetry, wherein a target concentration-independent or less concentration-dependent signal is generated at a particular frequency or frequencies. Exemplary voltammetry includes cyclic voltammetry, differential pulse voltammetry, alternating current voltammetry, potentiometric method, or amperometric method. An appropriate excitation waveform can be selected as is known in the art. In addition to square waves, other waveforms may be used, such as linear scan, triangular, differential pulse, and step waveforms.

The excitation pulse frequency can be selected for effective signal generation. For example, a voltage step in the range of +/- 0.1V to 0.5V has a repetition rate of 1 to 10,000Hz, for example about 5Hz, 10Hz, 20Hz, 30Hz, 40Hz, 50Hz, 60Hz, 70Hz, 80Hz, 90Hz, 100Hz, 110Hz, 120Hz, 130Hz, 140Hz, 150Hz, 160Hz, 170Hz, 190Hz, 200Hz, 250Hz, 300Hz, 350Hz, 400Hz, 450Hz, 500Hz, 600Hz, 700Hz, 800Hz, 900Hz, 1000Hz, 1500Hz, 2000Hz, 3000Hz, 4000Hz, 5000Hz and 10,000 Hz, and any intermediate frequency between 1 and 10,000 Hz.

Acquisition of time-resolved current measurements on a time scale of microseconds to milliseconds will generally be performed. Typical transients have a duration in the range of 10 to 100 ms and a selected time interval, e.g., approximately 1 μs, 2 μs, 3 μs, 5 μs or 10 μs, 50 μs, 60 μs, 70 μs, 80 μs, 90 μs, 100 μs, 200 μs, 300 μs, 10, 400 μs , 500 μs, 600 μs, 700 μs, 800 μs, 900 μs, 1 ms, 5 ms, 10 ms, 20 ms, 30 ms, 40 ms, 50 ms, 60 ms, 70 ms, 80 ms, 90 ms, or sampling at intervals greater than 100 ms.

In some sensor systems, the interconversion kinetics between the coupled and uncoupled states of a recognition element is faster than the electron transfer event measured by the sensor. Thus, the observed current transients reflect the population weighted average of the coupled and uncoupled states.

Sensor placement and operating conditions. Sensors are utilized in selected operating conditions that include various aspects of the sensing process. The operating conditions may include any combination of factors that affect the operation and output of the sensor.

In a first aspect, the operating conditions comprise the type of sample to be analyzed. A target species of the invention is evaluated in a sample. The sample will contain a liquid. A sample may be whole blood, serum, saliva, urine, sweat, interstitial fluid, spinal fluid, cerebral fluid, tissue exudate, broken tissue sample, cell solution, intracellular compartment, water, wash water, wastewater, groundwater, food, beverage, or other biological and environmental It may contain enemy samples. In some embodiments, the sample is from a subject, eg, a human patient or a non-human animal, such as a veterinary subject or test animal. In one embodiment, the sample comprises flowing whole blood, ie, blood sampled by a sensing system comprising a sensor implanted in a living body (eg, circulatory system) of a subject.

In one embodiment, the sample is treated prior to measurement. Examples of treatment include filtration, dilution, buffering, centrifugation, and application of other materials or processes to the sample prior to analysis. In some embodiments, the sample is not treated prior to performing the measurement, eg, the sample is undiluted, undiluted, undiluted, eg, whole blood.

In a second aspect, the operating conditions comprise assay conditions. Typical analysis conditions will represent reaction conditions for analysis, such as sample volume, temperature, pH, and the like.

The calibration-free and drift-free method of the present invention is particularly suitable for in vivo measurements. In one embodiment, the electrochemical sensor is configured for placement of a sensing element in vivo. In one embodiment, the sensing element comprises microwires. In various embodiments, the sensing element or housing of the sensing system may be inserted, implanted, or disposed within the body of a living organism. The sensing element of the sensing system may be implanted in the circulatory system, subcutaneously, intraperitoneally, intratracheally, or in another body compartment, wherein the sensing element is placed in an in vivo fluid, eg, interstitial fluid, blood, eg, flowing whole blood. exposed An implanted sensing system of the present invention may include an implanted sensing element coupled to a component external to the body (eg, by lead, wire, or wireless communication means), wherein the external component generates pulses, data Acquire or process. Alternatively, one or more auxiliary components to the sensing element, or even the entire sensing system of the present invention, may be configured to communicate (eg, by lead, wire, or wireless communication device) to an external device for data collection via the main body. can be transplanted to

In one embodiment, the methods of the present invention are performed in a feedback controlled dosing system, as is known in the art, wherein the drug is administered to maintain blood concentrations within a therapeutic or safety index. For example, in such methods and systems, the methods of the present invention are applied using an electrochemical sensing element implanted in a subject to measure the concentration of a target species, wherein the target species is a drug, a metabolite of the drug, or the drug It is a biomarker indicating that it should be administered. When the detected level of the target species indicates that the subject is in need of administration of the drug or other agent, the implanted pump or other agent delivery means can be used to maintain the concentration of the drug or agent within the desired range of the drug metered in dose. or administering the agent.

In another context, the sensors of the present invention are used for long-term and/or continuous monitoring of the environment or industrial sites, for example rivers, seas, water treatment plants, industrial facilities, food processing facilities, etc.

The sensors of the present invention may also be used in ex vivo and diagnostic applications. In one embodiment, a method of the invention comprises taking a sample from a living organism, and determining the concentration of a target species in the sample by an uncalibrated measurement. In one embodiment, the sensor of the present invention is used in a point-of-care test system. For example, in one embodiment, the sample is a blood sample, eg, an autologous pin or finger prick blood sample, or a urine, sweat or saliva sample. In such embodiments, the electrochemical sensor may be disposed in a housing such as a well, slide, lab-on-chip, microfluidic chamber, or other device.

drift compensation. In one aspect, the scope of the present invention includes a method of measuring the concentration of a target species in a sample using a novel electrochemical sensor, wherein drift correction is used to account for signal drift. The secondary moiety of the electrochemical sensor provides a means to implement drift compensation by imparting or enhancing the frequency dependence of the sensor output.

In various embodiments, the scope of the present invention includes a method for drift-corrected measurement of a concentration of a target species in a sample using an electrochemical sensor, the method comprising the steps of:

positioning the sensing element of the electrochemical sensor to be exposed to the sample;

wherein the sensing element comprises an electrode; the electrode is functionalized with a plurality of recognition elements, the recognition elements are capable of selectively binding to a target species, the recognition elements are functionalized with one or more redox reporters; The electrode is also functionalized with a plurality of charge transfer-modulating moieties, wherein the charge-transfer control moiety does not substantially bind to the target species;

applying a series of excitation pulses to the sensing element at a selected unresponsive frequency to generate a reference signal, wherein the unresponsive frequency is such that the signal is not measurably affected by the presence of the target and is minimally responsive to, or It is the frequency that responds negatively to the presence of a target;

applying a series of excitation pulses to the electrochemical sensor at a selected response frequency to generate a measurement signal, wherein the response frequency is a frequency at which the signal depends on a target concentration in the sample;

subtracting the reference signal from the measurement signal to determine a drift corrected signal value; and

determining the concentration of the target in the sample by applying a mathematical relationship between the drift-corrected signal value and the target concentration to the measured drift-corrected signal.

The first non-response frequency and the second response frequency are specific sensors, or specific sensor classes (ie, sensors of the same design and configuration, sensors manufactured from a specific lot, or other groups of sensors expected to have similar behavior) in a selected set of operating conditions. can be determined by one of ordinary skill in the art by routine experimentation with , wherein the sensor output is measured over a range of frequencies and over a range of target concentrations to determine the frequency at which the signal output is independent of the target concentration and the frequency at which the output responds to the target concentration. In some implementations, the non-response frequency is a relatively lower frequency and the response frequency is a relatively higher frequency. In various embodiments, a non-response frequency may include a frequency at which a signal does not measurably respond to a target concentration. In some embodiments, the non-response frequency is the frequency at which the signal least responds to the target concentration, ie the frequency at which it is least responsive over a range of frequencies. In some embodiments, the no-response frequency is the frequency at which the sensor output responds negatively to the target concentration.

The mathematical relationship between the drift-corrected output and the target concentration can be determined by one of ordinary skill in the art by routine experimentation with a particular sensor or particular class of sensors, under a selected set of operating conditions, where the sensor output is the difference between the drift-corrected sensor output and the sample. Measurements are made over a range of target concentrations at response and non-response frequencies to determine a mathematical relationship between target concentrations. For example, a standard curve, transform coefficient, or other mathematical relationship may be generated that relates a drift corrected normalized signal to a target concentration. For example, a drift corrected normalized signal versus concentration plot as in FIG. 4C can be used to relate the drift corrected signal to a target concentration. For example, in one embodiment, a calibration curve can be generated by exposing the sensor to a first calibration sample, e.g., a zero target, followed by exposing the sensor to a series of calibration standards at different target concentrations, and monitoring the sensing and reference currents. It is created by generating a standard curve over the dynamic range of the sensor.

Uncalibrated measurement. In some situations, for example, for sensors deployed in vivo, or in other situations where sensor calibration is impractical or burdensome, calibration-free measurements are desirable. The novel sensor of the present invention enables uncorrected measurements by providing a reference signal that is independent of the target concentration and can be used in ratiometric methods to normalize the signal output.

In various embodiments, the scope of the present invention includes a method for determining the concentration of a target species in a sample using an electrochemical sensor comprising the steps of:

acquiring a measurement value by means of a sensing element, wherein the sensing element comprises an electrode; the electrode is functionalized with a plurality of recognition elements, the recognition elements are capable of selectively binding to a target species, the recognition elements are functionalized with one or more redox reporters; The electrode is functionalized with a plurality of charge transfer-regulating moieties,

obtaining a baseline reference signal by exposing the sensing element to a calibration sample of a known target concentration and (1) applying a series of excitation pulses to the sensing element at a selected non-response frequency, wherein the non-response frequency is at which the signal is measurable by the presence of the target. is the frequency that is unaffected, minimally responsive to the presence of the target, or responds negatively to the presence of the target; and (2) obtaining a reference measurement signal by applying a series of excitation pulses to the sensing element at a selected response frequency, wherein the response frequency is a frequency at which the signal depends on the target concentration in the sample;

exposing the sensing element to a sample of unknown target concentration (1) obtaining a reference signal by applying a series of excitation pulses to the sensing element at a non-responsive frequency; and (2) obtaining a measurement signal by applying a series of excitation pulses to the sensing element at a response frequency;

calculating a normalized reference signal that is a ratio of a reference signal obtained from a sample of unknown target concentration to a baseline reference signal measured in a sample of known target concentration;

calculating a normalized measurement signal that is a ratio of the measurement signal obtained in the sample of unknown target concentration to the baseline reference signal obtained in the sample of known target concentration;

obtaining a normalized drift-corrected signal by subtracting a normalized reference signal representing drift from the normalized measurement signal; and

determining a target concentration of the sample by applying a mathematical relationship between the normalized drift-corrected signal and the target concentration to the measurements of the normalized drift-corrected signal.

The reference calibration sample may be a sample of known target concentration. In one embodiment, the reference sample comprises a blank, ie a sample in which zero target is present, or in an alternative embodiment it may be a reference sample containing a known concentration of target.

It will be understood that the order of the measurement steps is not critical to the method, wherein the reference and sense signals may be acquired in any order and the reference and measurement signals may be acquired in any order. In some embodiments, measurements of the reference reference and sense signals are performed prior to placing the sensor on the sample. For example, in one embodiment, the reference sensing and measurement of the reference signal are performed in a calibration solution and the sensor is then placed in vivo for collection of the sensing and reference measurement signal.

In one embodiment, the normalized drift corrected signal is determined as follows:

[Equation 1]

where S _cor is the normalized drift-corrected signal, i _S is the measured signal measured in a sample of unknown concentration; i _R is the reference signal measured in a sample of unknown concentration; i _S0 is the baseline sensing signal measured in a calibration sample of known concentration. i _R0 is the baseline reference signal measured in a calibration sample of known concentration.

The normalized drift-corrected change in the signal is converted to a target concentration value by comparing it to a standard curve, applying a transformation factor, or performing any other step of correlating the measured normalized drift-corrected signal to the target concentration. can be

An improved camptothecin aptamer. In one embodiment, the scope of the present invention includes novel camptothecin binding aptamers. The camptothecin aptamers described in Fujita, et al., 2013. Efficacy of Base-Modification on Target Binding of Small Molecule DNA Aptamers, Pharmaceuticals 6: 1082-1093 are known in the art. A 40-base cleavage of this aptamer, referred to herein as C40, comprises SEQ ID NO: 1, which is a target-recognizing G-quadruplex flanked by a 12-base-pair stem as shown in FIG. 2A . Folds. The scope of the present invention includes modifications of the C40 aptamer.

In one embodiment, the camptothecin aptamer is CA40_1MM comprising SEQ ID NO: 2 (5'-ACGCT CCGGACTTGG GTGGG TGGGT TGGGG TACGG TGTGT-3'), wherein nucleotides 1-13 and 29-40 constitute the stem portion and Nucleotides 14-28 constitute the G-quartet binding pocket.

In one embodiment, the camptothecin aptamer is CA40_2MM comprising SEQ ID NO: 3 (5'-ACGCT CTGGA CTTGG GTGGG TGGGT TGGGG TACGG TGTGT-3'), wherein nucleotides 1-13 and 29-40 constitute a stem portion and nucleotides 14-28 constitute the G-quartet binding pocket.

In one embodiment, the camptothecin aptamer is CA36 comprising SEQ ID NO: 4 (5'-GTCCC GGACT TGGGT GGGTG GGTTG GGGTA CGGTG C-3'), wherein nucleotides 1-11 and 28-36 constitute the stem portion. and nucleotides 12-27 constitute the G-quartet binding pocket.

In one embodiment, the camptothecin aptamer is CA32 comprising SEQ ID NO: 5 (5'-TCCGG ACTTG GGTGG GTGGG TTGGG GTACG GT-3'), wherein nucleotides 1-9 and 26-32 constitute the stem portion and Nucleotides 1-25 constitute the G-quartet binding pocket.

In one embodiment, the camptothecin aptamer is CA28 comprising SEQ ID NO: 6: (5'-CGGAC TTGGG TGGGT GGGTT GGGGT ACG-3'), wherein nucleotides 1-7 and 24-28 constitute the stem portion and Nucleotides 8-23 constitute the G-quartet binding pocket.

In one embodiment, the camptothecin aptamer is CA16 comprising SEQ ID NO: 7 (5′-GGGTGGGTGGGTTGGG-3′) comprising only a G-quadrued binding pocket.

In one embodiment, the scope of the invention is C40 (SEQ ID NO: 1); CA40_1MM, (SEQ ID NO: 2); CA40_2MM (SEQ ID NO: 3); CA36 (SEQ ID NO: 4); CA32 (SEQ ID NO: 5); CA28 SEQ ID NO: 6); and a sensor for the measurement of camptothecin and derivatives thereof comprising a recognition element comprising a camptothecin aptamer selected from the group consisting of CA16 (SEQ ID NO: 7).

In one embodiment, the scope of the invention is C40 (SEQ ID NO: 1); CA40_1MM, (SEQ ID NO: 2); CA40_2MM (SEQ ID NO: 3); CA36 (SEQ ID NO: 4); CA32 (SEQ ID NO: 5); CA28 SEQ ID NO: 6); and a method of determining the concentration of a camptothecin compound in a sample comprising the use of an electrochemical sensor comprising a recognition element comprising a camptothecin aptamer selected from the group consisting of CA16 (SEQ ID NO: 7). The camptothecin compound may be camptothecin or a derivative thereof, for example irinotecan, topotecan, belotecan or deruxtecan.

Example.

Example 1. Measurement of in situ hyperdissolution pharmacokinetics of chemotherapeutic agent irinotecan in vivo.

Introduction. The goal of personalized medicine is to provide treatment that is precisely tailored to the individual. To this end, the ability to measure drugs in vivo with super-resolution allows clinicians to make drug dosage decisions based on high-precision, patient-specific pharmacokinetic measurements rather than indirect predictors of drug metabolism such as age, body mass, or pharmacogenetics. can do. Ultimately, the ability to measure drugs in the body in real time will enable closed-loop feedback controlled delivery, significantly improving dosing accuracy by actively responding to minute-to-minute fluctuations in a patient's metabolism. However, the development of these technologies faces significant obstacles. First, in vivo sensors must be small enough to be mounted without causing undue damage to the body. Second, it cannot require the addition of exogenous reagents or the use of batches such as washing or separation. Third, it should be measured at a faster frequency than the pharmacokinetics of the drug. Finally, it must be selective and stable enough to allow long-term operation in the complex and fluctuating environments found in vivo. For this purpose, an improved electrochemical aptamer based (E-AB) sensor is described herein.

The E-AB sensor uses an electrode-bound, redox reporter-modified aptamer as a recognition element (Fig. 1a). Binding of the target molecule to this aptamer induces a conformational change that produces an easily measured electrochemical output (eg, square wave voltammetry) without the need for reagent addition or washing steps. Because E-AB signaling is generated by binding-induced conformational changes, unlike the case of most other drug-free biosensor architectures, by target adsorption to the sensor surface, the E-AB sensor is largely insensitive to non-specific adsorption and in vitro. In addition, it supports the measurement of several hours of biological fluids in vivo. Finally, since their signaling occurs due to target binding alone, not from the chemical reactivity of the target, but not in the case of continuous glucose monitors, the E-AB sensor uses amino acids, two of the analytes measured in vivo. It is a platform technology that can be generalized to a wide variety of analytes, including glycosides and doxorubicin. Based on this basis, for the camptothecin class of anticancer agents, an important class of chemotherapeutic agents used in the treatment of a range of human cancers, the preparation and characterization of E-AB sensors suitable for in situ measurements of the body are described herein. do.

Results and Discussion. As a recognition element of the sensor, a DNA aptamer that binds camptothecin may be used. Specifically, the 40-base version of the aptamer, termed CA40, which folds into a target recognition G-quartet flanked by a 12-base pair stem (Fig. 2a), has a dissociation constant of about 475 nM when the unmodified aptamer is not in solution. Binds to the tothecin, irinotecan (Fig. 2b). To apply this to the E-AB sensor, its 3' end was modified with a methylene blue redox reporter and deposited on a gold electrode via a 6 carbon thiol at its 5' end (Fig. 1a). Upon electrochemical interrogation of the buffer-generated sensor, the expected Langmuir isothermal binding resulted in an estimated KD of 126+/-24 mM, a signal gain of 84 of 4% (relative change in signal upon addition of a saturated target) (Fig. 2c) and It was observed with association and dissociation kinetics too fast to measure (time constant < 5 s at clinically relevant concentrations, Figure 2d). The lower affinity exhibited by the aptamer with respect to the sensor may arise from interactions with the electrode surface, as it is known to destabilize the folding (and thus hindering binding) surface-attached oligonucleotides. Nevertheless, when tested in buffer, the sensor supports the detection of irinotecan in the clinically relevant 1–15 mM range. However, when tested in whole blood, the (apparent) affinity of the surface-bound aptamer is still worse (K _D =291+/-15 mM), probably because the concentration of free drug decreases due to protein binding. To make matters worse, under these conditions the gain drops to about 15%, leaving the useful dynamic range outside the clinically relevant concentration window (Fig. 2c).

To improve sensor performance in whole blood, camptothecin-binding aptamers were redesigned to better settle in an "unfolded" state in the absence of a target to increase the sensor's gain. To this end, the double-stranded stem of the aptamer was destabilized through the introduction of cleavage (CA36, CA32, CA28 CA16) or one (CA40_2MM) or two (CA40_2MM) GT mismatches ( FIG. 3A ). Nucleic acid fold predictor NUPACK software was used to estimate. This strategy reduces the stability of the folded aptamer from 31.8 kJ mol ⁻¹ of the C40 parent to 0.3 kJ mol ⁻¹ of the CA28. Characterizing the sensor fabricated using this modification (Fig. 3a), dissociation constants were obtained in the range of 38 11 mM to 254 48 mM when tested in working buffer, with a signal gain of up to 755% (Fig. 3b). In whole blood, the test ( FIG. 3c ) reduces both gain and apparent affinity once again. However, even under these more difficult conditions, sensors using CA32, CA28 and CA16 variants supported high-gain E-AB detection.

After achieving good in vitro performance with the sensor using the CA32 variant, the sensor was adapted for field use in live rat veins. Under these conditions, E-AB sensors often exhibit significant baseline drift. Previously, this was corrected using a “dynamic differential measurement” approach that exploits the generally strong square-wave frequency dependence of the E-AB signal gain. In particular, the signal gain of the previously described E-AB sensor is so large that it exhibits a “signal-on” (targeted coupling increases the signal current) response at some square wave frequencies and gains or even “unobservable” at other square wave frequencies. It exhibits "signal-off" behavior. Conveniently, the signals obtained in these different regimes are drifted together and taking their difference via KDM eliminates the drift seen in vivo. And since the two signaling currents react in opposite directions when the target is present, taking the difference improves the signal gain.

In contrast, the gain of camptothecin-detecting E-AB sensors (CA32, CA28, and CA16) is only a weak function of the square wave frequency, so the development of a new approach for performing KDM is required. To improve the frequency dependence of the sensor gain supporting KDM drift correction, a second reporter-modified DNA strand was added to the sensor that (1) transfers electrons faster than the aptamer and (2) does not respond to the presence of the target. . The reason for doing so is that the gain of the final sensor will be low at frequencies where this "unresponsive" DNA dominates the signal (ie at higher frequencies), and will instead be higher at frequencies where the aptamer dominates the signal gain. To achieve this, the CA32 aptamer modification consists of a secondary moiety comprising an unstructured 10-base strand (SEQ ID NO: 8) consisting of a random sequence of adenine and thymine, which is known to transfer electrons at a rate of 80 per second. deposited together. As expected, a sensor using a 1:1 mixture of this sequence and an aptamer achieves a frequency dependent gain sufficient to enable KDM (Fig. 4b). Surprisingly, the final frequency dependence is so strong that the sensor's gain becomes slightly negative at low frequencies, a record that does not match the expectations described above (if current is added, the gain cannot drop below zero). This is presumed to occur due to the interaction between the two sequences on the surface, which alters the electron transport kinetics. However, regardless of their origin, this effect supports accurate KDM drift correction. Specifically, to perform KDM, using signals obtained at 10 Hz (signal-off) and 120 Hz (signal-on), irinotecan was administered in total 0.5 to 15 mM (0.06 to 10 mg/mL) buffer and whole blood over the human therapeutic range. All can be monitored (Fig. 4c).

The KDM-calibrated intrinsic E-AB sensor readily supports real-time, high-frequency irinotecan measurements in situ in the body of live mice. A sensor using gold, platinum, and silver wires with a diameter of 75 mm as a working electrode, a counter electrode, and a reference electrode, respectively, was fabricated ( FIG. 1B ). The sensor was placed in the jugular vein of anesthetized Sprague-Dawley rats via a previously placed 22 gauge catheter ( FIG. 1C ). Testing this with a single intravenous injection of irinotecan (20 mg/kg) resulted in the observed signals at 10 and 120 Hz responding to the drug, but also accompanied by the expected signal drift. The gain observed at 10 Hz becomes slightly positive under these conditions. KDM could be used to correct for drift in the sensor and restore a stable baseline, thus allowing continuous real-time measurements of drugs at treatment-relevant concentrations.

To further characterize the performance of the camptothecin sensor, this sensor was used to monitor sequential intravenous injections of irinotecan (10 and 20 mg/kg). The resulting maximum concentrations (C _MAX 1/4 39.8 3.2 mM and 20.9 +/- 2.0 mM, respectively; excitation and below confidence intervals reflect standard errors calculated from standard fit) and distribution ratios (a 1/4 0.58 +/-0.07 min and 0.48 min) are similar to those seen in previous studies using in vitro drug level measurements. In contrast, the observed clearance rates (b 1/4 9.2 +/- 1.4 min and 8.4 +/- 2.2 min) are faster than previously reported, resulting in reduced "area under the curve" for the drug. This discrepancy is thought to have arisen because previous work has used chromatographic and mass spectrometry methods to measure total drug levels (requires removal of blood samples from the animal's body and extraction of the entire drug into a buffer). In contrast, E-AB sensors measure free drug, which is the proportion of pharmacologically active drug. And, in general, the release and clearance of free drug is faster than the total drug because drugs that strongly interact with plasma proteins tend to be eliminated more slowly than drugs that do not.

E-AB-induced measurement of irinotecan pharmacokinetics represents a significant advance over previous pharmacokinetic studies of camptothecin. For example, the 20 second time resolution of our measurements (defined as the time required to perform the two square wave scans required to perform KDM) is at least 10 times better than the highest time resolution previous studies. Moreover, all previous studies reported mean plasma level measurements for multiple animals, eliminating the ability to explore inter-subject pharmacokinetic variability. In contrast, current E-AB-derived measurement parameters provide 300 time points per hour in each animal, thus determining individual pharmacokinetics with great precision. Since the excretion phase of irinotecan exhibits significant inter-patient variability (due to drug-drug interactions, variability in health status, and pharmacogenetics), the latter point may be of clinical significance. To demonstrate the ability to measure this variability, we sequentially injected 10 mg/kg and 20 mg/kg of irinotecan to rats ( FIG. 7 , representative data from a single rat). Outcome measures show only small (10-20%) variability in the percentage of C _MAX or distribution steps. However, in contrast, drug excretion rates and clearance values varied several-fold from rat to rat. Since all animals used in this experiment were healthy male mice, these pharmacokinetic differences were only due to metabolic variability between animals.

The excretion rate and clearance rate of irinotecan are pharmacokinetic parameters used to determine its optimized and personalized dosage during chemotherapy. To measure these parameters more accurately, we administered a large amount of drug (60 mg/kg) over a long period of time. The resulting higher plasma concentration results in a longer measurement period (after delivery has ceased) before reaching the sensor's detection limit (Figures 5a and 5b). Thus, the 150 measurements achieved in a single pharmacokinetic profile are significantly less than the two measurements used for typical “peak and trough” clinical measurements, and significantly more than those generated in conventional in vitro studies, typically obtaining less than 12 measurements per profile. An estimate of the exact drug clearance half-life (10.4 +/- 0.4 min) and clearance (18.6 +/- 1.4 mL per minute) was generated.

The following describes an intrinsic E-AB sensor that supports the super-resolution measurement of the in situ anticancer drug irinotecan in vivo for several hours. The sensor design required redesigning the parent aptamer to support high-gain E-AB signals and development of a new method to ensure sufficient frequency-dependent signal gain to support KDM-based drift compensation. Plasma irinotecan levels were measured with micromolar resolution and second time resolution using the final sensor, the latter being improved by a factor of 10 compared to the previous study. The final measure defines the pharmacokinetics of irinotecan in individual animals, providing an unprecedented high-resolution review of the drug's intersubject pharmacokinetic variability.

The E-AB sensor is a platform technology that supports high-frequency real-time measurement of specific molecules (regardless of chemical reactivity) in situ in vivo. This versatility, when combined with the convenience and precision of the platform, offers significant opportunities to improve drug administration. For example, as noted above, irinotecan undergoes severe patient-to-patient metabolic variability, resulting in toxicity and increased side effects. However, because current methods of measuring plasma drug levels are slow and cumbersome, FDA has used pharmacological estimates of metabolism as the primary means of reducing the risks associated with this variability. In this regard, the ease with which E-AB sensors provide high-precision, patient-specific measurements (as opposed to indirect estimates) of drug clearance suggest that the platform may provide a valuable adjunct to chemotherapeutic treatment.

In addition to improving the precision and accuracy of personalized dose determination, E-AB-guided measurements may support a new paradigm for personalized drug delivery. In particular, real-time concentration information provided by the E-AB sensor can be used to inform closed-loop feedback controlled drug delivery. Here, the administered drug rate can be optimized multiple times per minute to maintain plasma drug concentrations at predefined values with greater than 20% precision despite 3-fold hourly changes in pharmacokinetics. This approach to drug delivery provides an unprecedented means of overcoming pharmacokinetic variability and improving the overall efficacy and safety of treatment. Given this, with the improvement of the present invention, the drug-detecting E-AB sensor provides a powerful new tool in the clinician's arsenal.

Materials and Methods

sensor fabrication. A catheter (22G) and a 1 mL syringe were used. Polytetrafluoroethylene-insulated gold, platinum, and silver wires (75 μm in diameter) were used. To use a silver wire as a reference electrode, it was immersed in concentrated sodium hypochlorite (commercial bleach) overnight to form a silver chloride film. The wires were electrically insulated using heat-shrinkable polytetrafluoroethylene insulation. A custom, open, mesh-covered 3-channel connector cable was used to fabricate the in vivo probe.

oligonucleotides. For sensor fabrication, an aptamer modification modified with methylene blue attached to a thiol-C ₆ -SS group at the 3' end and an amine at the 5' end by a 6 carbon linker was purchased. Oligonucleotides were dissolved in a buffer (100 mM Tris buffer, 10 mM MgCl ₂ , pH 7.8) at a concentration of 100 μM, aliquoted and stored at -20°C.

electrode polishing and cleaning. E-AB sensors used in vitro are described, for example, in B. Sanavio and S. Krol, Front. Bioeng. Biotechnol., 2015, 3, 20 were fabricated using established methods as described. Briefly, E-AB sensors were fabricated on rod gold disk electrodes. Disc electrodes were prepared by wetting with a 1 μm diamond suspension slurry and then grinding on a microcloth pad moistened with an aqueous suspension of 0.05 μm alumina powder. For each polishing step, the electrode was sonicated in a 1:1 water/ethanol solution for 5 min. The electrodes were then electrochemically cleaned according to the following procedure: a) the electrodes were placed in a 0.5 M NaOH solution and a voltage of 2 V per second was 1000-2000 scans were performed at scan rate; b) The electrode was transferred to a 0.5MH ₂ SO ₄ solution and an oxidation potential of 2V was applied for 5 seconds using chronoamperometry. After applying a reducing potential of -0.35 V for 10 s c) using cyclic voltammetry, the electrode was rapidly cycled between -0.35 and 1.5 V (4 V per sec) for 10 scans in the same solution in the same solution, then using the same potential window per second Two cycles recorded at 0.1V were recorded.

Electrode functionalization. For the classical monolayer functionalized E-AB sensor consisting of a single DNA probe, the DNA probe was first reduced by treatment for 1 h in a 10 mM tris(2-carboxyethyl)-phosphine hydrochloride (TCEP) solution at room temperature in the dark. (100 μM). It was then dissolved in "assembly buffer" (10 mM Na ₂ HPO ₄ containing 1M NaCl and 1 mM MgCl ₂ at pH 7.3) to a final concentration of 500 nM. Then, the electrochemically cleaned gold electrodes were immersed in 200 μL of this solution in the dark for 1 hour. The electrode surfaces were then rinsed with distilled water and incubated overnight at 4° C. in an assembling buffer containing 5 mM 6-mercaptohexanol, followed by further rinsing with distilled water before use. For the E-AB sensor functionalized with a mixed monolayer formed by a CA32 aptamer and a linear probe, both DNA probes were first incubated for 1 h in 10 mM tris(2-carboxyethyl)-phosphine hydrochloride (TCEP) solution at room temperature in the dark. It was reduced by treatment (100 μM). They were then dissolved in buffer (10 mM Na ₂ HPO ₄ containing 1 M NaCl and 1 mM MgCl ₂ at pH 7.3) using two DNA probes in different ratios. The total oligonucleotide concentration was kept constant at 500 nM to maintain a constant DNA probe packing density. The electrochemically cleaned gold electrodes were then immersed in 200 µL of this solution consisting of two DNA probes in the dark for 1 hour. After this step, the electrode surface was rinsed with distilled water and incubated overnight at 4° C. in an assembling buffer containing 5 mM 6-mercaptohexanol, followed by further rinsing with distilled water before use.

In vivo sensors: electrode fabrication. E-AB sensors used in vivo were fabricated as described in a previous report (Tucker, Pharm. Res., 2017, 34, 1539-1543, N. Arroyo-Curras, G. Ortega, DA Copp, KL Ploense, ZA). Plaxco, TE Kippin, JP Hespanha and KW Plaxco, ACS Pharmacol. Transl. Sci., 2018, 1, 110-118). The sensors were made by cutting pieces of pure gold (7.75 cm long), platinum (7.25 cm), and silver (6.75 cm) wires. About 1 cm of insulation was removed from both ends of this wire using a surgical blade to allow electrical contact. Then each of these is soldered to one of the three ends of the connector cable using 60% tin/40% lead rosin-core solder (0.8mm in diameter), then the wire, except for a small window about 5mm on the edge of each wire, Heat was applied to the shrink tubing around the body to attach them together. The wires were connected in a layered manner, and first only the gold wire was insulated, then the gold wire and the platinum wire were insulated together, and finally all three wires were insulated together. The purpose of this three-layer thick insulator was to impart mechanical strength to the malleable probe body. To prevent electrical shorts between the wires, different lengths were used for each wire as previously described. A sensor window (ie, an area without insulation) of gold wire was cut to a length of about 3 mm. To increase the surface area of the gold working electrode (to obtain a larger peak current), the sensor surface was immersed in 0.5 M sulfuric acid and the potential of E _initial = 0.0 V to E _high = 2.0 V for Ag/AgCl was applied back and forth for 16000 pulses. It was electrochemically roughened by increasing. Each potential step was 20 ms duration with no latency between pulses.

Electrode functionalization. The E-AB sensor used in vivo was functionalized with a mixed monolayer formed by the CA32 aptamer and the linear probe SEQ ID NO:8. Two DNA probes (100 μM) were reduced by treatment (100 μM) in a solution of 10 mM tris(2-carboxyethyl)-phosphine hydrochloride (TCEP) in the dark at room temperature for 1 hour. This was then made into an "assembly buffer" (containing 1M NaCl and 1 mM MgCl at pH 7.3) using a 1:1 mixture of _two DNA probes (i.e., 250 nM for each DNA probe) for a total oligonucleotide concentration of 500 nM. 10 mM Na ₂ HPO ₄ ). The electrochemically roughened gold electrode was rinsed with deionized water and immersed in 200 μL of this solution consisting of two DNA probes for 1 hour in the dark. The electrode surfaces were then rinsed with distilled water and incubated overnight at 25°C in an assembling buffer containing 5 mM 6-mercaptohexanol, followed by further rinsing with distilled water before use.

In vitro measurements. Electrochemical measurements were performed at room temperature using a CHI660D potentiometer with a CHI684 multiplexer (CH Instruments, Austin, TX) and a standard three-electrode cell containing a platinum counter electrode and an Ag/AgCl (3M KCl) reference electrode. Square wave voltammetry (SWV) was performed using a potential window of -0.1 to -0.4 V, a potential step of 1 mV and an amplitude of 50 mV.

Titration curve. Experimental titration curves using three E-AB sensors modified with oligonucleotide probes and using SWV frequencies of 10 and 120 Hz, experimental titration curves were obtained using 10 mL of working buffer (137 mM NaCl, 2.7 mM KCl, _2.7 mM Na2HPO at pH 7.3 _{4 10} mM, KH ₂ PO ₄ 1.8 mM). Preliminary treatment was initially performed in the absence of irinotecan by interrogating the electrode with 30-60 scans until a stable current peak was obtained. When the signal of the sensor was stabilized, the concentration of irinotecan was increased and the sensor was irradiated after 10 minutes. The electrochemical signal (peak current) of each sensor was plotted as a function of irinotecan concentration and then fitted using the single site binding mechanism equation:

Equation 2:

where [Irinotecan] is the irinotecan concentration, C ^{[Irinotecan raw]} is the signal current in the presence of different concentrations of irinotecan, C ^{0 raw} is the background current in the absence of irinotecan, and C ^[Irinotecan] is the irinotecan at saturation concentration This is the current signal and KD is the dissociation constant of the surface-associated aptamer. Using the values estimated from the previous fitting procedure, the raw signal current was converted to % signal change (C%) using the following equation:

Equation 3:

Binding curves performed on whole blood were performed according to the previously described experimental approach using a closed loop system with a continuous flow of whole blood through a circulatory pump (total volume of 20 mL at 1 mL/s).

sensor equilibration time. The equilibration time of the sensor was determined using the experimental approach above and interrogating the sensor every 5 s in the working buffer using a SWV frequency of 500 Hz. After a stable current baseline (5 min) was reached, different concentrations of irinotecan were added to the solutions (1 μM to 1 mM) and the voltammetry signal was monitored for at least 5 min. The sensor was then transferred to a simple working buffer without irinotecan and voltammetric signals were collected over 5 min. The observed signal change was fitted to a single exponential decay to obtain the sensor's equilibrium time constant.

Square wave voltammetry frequency versus signal change (%) application. The dependence of the E-AB sensor signal was dependent on the change on the square-wave voltammetry frequency interrogating the sensor using various frequencies (5 Hz to 4000 Hz) in the absence and presence of saturating amounts of irinotecan (1 mM and 100 μM and an incubation time of 10 min). %). Peak currents collected in the absence and presence of a saturating amount of irinotecan (1 mM or 100 μM) were converted to % signal change using the above equation. The estimated signal change (%) was then expressed as a function of the relative square wave frequency.

In vivo measurements. All in vivo measurements were performed using a three-electrode setup in which the counter electrode was made of platinum wire and the reference electrode was a silver wire coated with a silver chloride film as described above. Measurements performed in vivo were recorded using a portable potentiometer of the CH instrument (Model 1242 B). Square wave voltammetry (SWV) was performed using a potential window of -0.1 to -0.45 V, a potential step of 0.001 V and an amplitude of 0.05 V.

For in vivo measurements, a sensor baseline was established 30 min prior to drug infusion. A 3 mL syringe filled with the target drug was connected to a sensorless catheter (located opposite the jugular vein where the sensor was mounted) and placed in an electric syringe pump. After establishing a stable baseline, the drug was infused through this catheter using 15 mg/mL irinotecan or 10 mg/mL 5-fluorouracil at a rate of 0.6 mL/min. After drug infusion, recordings were taken for up to 1 hour before the next infusion. Real-time plotting and analysis of voltage and current data was performed with the help of scripts written in Python.

animal. In vivo measurements were performed in male and female Sprague-Dawley rats (4-5 months old) weighing 300-500 g. All animals were housed in pairs in a standard light cycle room (08:00 on, 20:00 off) and had free access to food and water. For in vivo measurements, mice were induced under 5% isoflurane anesthesia in a Plexiglas anesthesia chamber. The mice were then maintained with 2-3% isoflurane gas for the duration of the experiment. A small incision was made on each vein and then each vein was isolated. A small hole was made in each vein with fine spring-loaded scissors. A silastic catheter inserted into the left jugular vein was fabricated with a bent steel cannula with a threaded connector and a silastic tube (11 cm, id 0.64 mm, od 1.19 mm). The EAB sensor was fabricated as previously described and inserted into the right jugular vein. Both the E-AB sensor and infusion line were tied in place with sterile 6-0 silk sutures and then tied with 30 units of heparin.

Pharmacokinetic analysis . Plasma irinotecan concentration-time curves were fitted to a bi-exponential equation describing a two-compartment model.

Equation 4: [irinotecan] = Ae ^-t/α + Be ^-t/β

where A and B are the amplitudes and α and β are the half-lives for distribution and elimination, respectively. To improve the accuracy of the fitting, a time span of 20 min was used for the pharmacokinetic curve obtained using an intravenous dose of 20 mg/kg of drug, and a time span of 20 min was used for the curve obtained using an intravenous dose of 10 mg/kg of irinotecan. A time span of 10 minutes was used for In addition, to improve the accuracy of the fitting equation, α and β were fitted for rat 1 and rat 3 to curves obtained by intravenous administration of 10 mg/kg of irinotecan. The peak concentration (C _MAX ) was obtained directly from the experimental value of each irinotecan concentration-time curve. The area under the curve (AUC) was calculated using the parameters A, B, α and β because integrating Equation 5 from 0 to infinity holds

Equation 5:

The removal rate (Cl _T ) was calculated by the following equation:

Equation 6:

The errors for C _MAX , AUC and Cl _T are increased from the errors related to the kinetic parameters A, B, α and β.

NUPACK simulation. NUPACK software (http://www.nupack.org/) was used to predict the folding free energy of the stem portion for CA40 aptamer modification. Aptamer sequences were analyzed by software using the following parameters. (a) temperature: 25° C.; (b) number of strand species: 1; (c) Maximum complex size: 4 (d) Oligo concentration = 500 nM; in advanced options; (e) [Na+] = 0.15M, [Mg++] = 0M; (f) Dangle treatment: some.

All patents, patent applications, and publications cited herein are hereby incorporated by reference to the same extent as if each independent patent application or publication were specifically and individually indicated to be incorporated by reference. The disclosed embodiments are presented for purposes of illustration and not limitation. While the invention has been described with reference to the described embodiments thereof, it will be understood by those skilled in the art that modifications may be made in the structure and elements of the invention without departing from the spirit and scope of the invention as a whole.

SEQUENCE LISTING <110> THE REGENTS OF THE UNIVERSITY OF CALIFORNIA <120> MODULATING ELECTRON TRANSFER KINETICS IN E-DNA-TYPE SENSORS <130> UCSB029PCT <140> <141> <150> 62/860,772 <151> 2019-06-12 <160> 13 <170> PatentIn version 3.5 <210> 1 <211> 40 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA40 aptamer, nucleotides 1-13 and 29-40: stem portion, nucleotides 14-28: G-quadruplex binding pocket <400> 1 acgctccgga cttgggtggg tgggttgggg tacggtgcgt 40 <210> 2 <211> 40 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA40_1MM aptamer, nucleotides 1-13 and 29-40: stem portion, nucleotides 14-28: G-quadruplex binding pocket <400> 2 acgctccgga cttgggtggg tgggttgggg tacggtgtgt 40 <210> 3 <211> 40 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA40_2MM aptamer, nucleotides 1-13 and 29-40: stem portion, nucleotides 14-28: G-quadruplex binding pocket <400> 3 acgctctgga cttgggtggg tgggttgggg tacggtgtgt 40 <210> 4 <211> 36 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA36 aptamer, nucleotides 1-11 and 28-36: stem portion, nucleotides 12-27: G-quadruplex binding pocket <400> 4 gctccggact tgggtgggtg ggttggggta cggtgc 36 <210> 5 <211> 32 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA32 aptamer, nucleotides 1-9 and 26-32: stem portion, nucleotides 1-25: G-quadruplex binding pocket <400> 5 tccggacttg ggtgggtggg ttggggtacg gt 32 <210> 6 <211> 28 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA28 aptamer, nucleotides 1-7 and 24-28: stem portion, nucleotides 8-23: G-quadruplex binding pocket <400> 6 cggacttggg tgggtgggtt ggggtacg 28 <210> 7 <211> 16 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> CA16 aptamer and sequence of the G-quadruplex binding pocket <400> 7 gggtgggtgg gttggg 16 <210> 8 <211> 10 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <220> <223> linear probe sequence <400> 8 attatttttt 10 <210> 9 <211> 5 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <400> 9 attat 5 <210> 10 <211> 15 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <400> 10 attatttttt attta 15 <210> 11 <211> 30 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <400> 11 attatttttt atttattttt attttatttt 30 <210> 12 <211> 40 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <400> 12 attatttttt atttattttt attttatttt attttttatt 40 <210> 13 <211> 60 <212> DNA <213> Artificial Sequence <220> <223> Description of Artificial Sequence: Synthetic oligonucleotide <400> 13 attatttttt atttattttt attttatttt attttttatt tatttttatt ttattttatt 60

Claims (21)

A sensing element for an electrochemical sensor for measuring a concentration of a target species in a sample comprising an electrode, comprising:
wherein the electrode is functionalized with a plurality of recognition elements, the recognition elements are capable of selectively binding to a target species, and the recognition elements are functionalized with one or more redox reporters; and
wherein the electrode is functionalized with a plurality of charge transfer-modulating moieties, wherein the charge transfer-modulating moieties do not substantially bind to a target species. The method of claim 1,
A sensing element, wherein the recognition element comprises a polynucleotide, polypeptide or chemical species. 3. The method of claim 2,
and the recognition element comprises a polynucleotide. 3. The method of claim 2,
wherein the polynucleotide comprises an aptamer. The method of claim 1,
wherein the charge transfer-regulating moiety comprises an oligonucleotide. 6. The method of claim 5,
A sensing element that is 5-20 nucleotides in length in an oligonucleotide. 6. The method of claim 5,
wherein the oligonucleotide consists of at least 90% thymine, adenine or a mixture of thymine and adenine. 6. The method of claim 5,
wherein the oligonucleotide is selected from the group consisting of SEQ ID NO: 8, SEQ ID NO: 9, SEQ ID NO: 10, SEQ ID NO: 11, SEQ ID NO: 12 and SEQ ID NO: 13. The method of claim 1,
wherein the charge transfer-regulating moiety is functionalized with one or more redox reporters. 7. The method of claim 6,
wherein the charge transfer-regulating moiety is not functionalized with a redox reporter. The method of claim 1,
The redox reporter of the recognition element and/or charge transfer-regulating moiety may be methylene blue, ferrocene, viologen, anthraquinone or any other quinone, ethidium bromide, daunomycin, organo-metal redox label, e.g. For example selected from the group consisting of porphyrin complexes or crown ether cycles or linear ethers, ruthenium, bis-pyridine, tris-pyridine, bis-imidazole, ethylenetetraacetic acid-metal complex, cytochrome c, plastocyanin and cytochrome c' The sensing element to be. The method of claim 1,
The sensing element is configured for in vivo use. A method for drift-corrected measurement of a concentration of a target species in a sample using an electrochemical sensor, the method comprising:
positioning the sensing element of the electrochemical sensor to be exposed to the sample; wherein the sensing element comprises the sensing element of any one of claims 1 to 12;
applying a series of excitation pulses to the sensing element at a selected unresponsive frequency to generate a reference signal, wherein the unresponsive frequency is such that the signal is not measurably affected by the presence of the target and is minimally responsive to or It is the frequency that responds negatively to the presence of a target;
applying a series of excitation pulses to the electrochemical sensor at a selected response frequency to generate a measurement signal, wherein the response frequency is a frequency at which the signal depends on a target concentration in the sample;
subtracting the reference signal from the measurement signal to determine a drift corrected signal value; and
determining the concentration of the target in the sample by applying a mathematical relationship between the drift-corrected signal value and the target concentration to the measured drift-corrected signal. 14. The method of claim 13,
wherein the sample is untreated. 14. The method of claim 13,
The sample is whole blood. 14. The method of claim 13,
The method of claim 1, wherein the sensing element is disposed within the living body. 13. A method of measuring a concentration of a target species in a sample using an electrochemical sensor, comprising the steps of: the electrochemical sensor comprising a sensing element comprising the sensing element of any one of claims 1-12 How to be:
obtaining a baseline reference signal by exposing the sensing element to a calibration sample of a known target concentration and (1) applying a series of excitation pulses to the sensing element at a selected non-response frequency, wherein the non-response frequency is at which the signal is measurable by the presence of the target. is the frequency that is unaffected, minimally responsive to the presence of the target, or responds negatively to the presence of the target; and (2) obtaining a reference measurement signal by applying a series of excitation pulses to the sensing element at a selected response frequency, wherein the response frequency is a frequency at which the signal depends on the target concentration in the sample;
exposing the sensing element to a sample of unknown target concentration (1) obtaining a reference signal by applying a series of excitation pulses to the sensing element at a non-responsive frequency; and (2) obtaining a measurement signal by applying a series of excitation pulses to the sensing element at a response frequency;
calculating a normalized reference signal that is a ratio of a reference signal obtained from a sample of unknown target concentration to a baseline reference signal measured in a sample of known target concentration;
calculating a normalized measurement signal that is a ratio of the measurement signal obtained in the sample of unknown target concentration to the baseline reference signal obtained in the sample of known target concentration;
obtaining a normalized drift-corrected signal by subtracting a normalized reference signal representing drift from the normalized measurement signal; and
determining a target concentration of the sample by applying a mathematical relationship between the normalized drift-corrected signal and the target concentration to the measurements of the normalized drift-corrected signal. 11. The method of claim 10,
wherein the sample is untreated. 11. The method of claim 10,
The sample is whole blood. 11. The method of claim 10,
The method of claim 1, wherein the sensing element is disposed within the living body. A camptothecin binding aptamer comprising a polynucleotide selected from the group consisting of SEQ ID NO: 2, SEQ ID NO: 3, SEQ ID NO: 4, SEQ ID NO: 5, SEQ ID NO: 6 and SEQ ID NO: 7 .

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